Systems for acquiring radiographic images are widely used in medical and dental care. Due to recent advances in component miniaturization, data transmission and processing speed, and with improved image processing and display capabilities, apparatus and methods for obtaining radiographic images directly in digital format are increasingly being used. With digital radiography, the radiation image exposures captured on radiation-sensitive phosphor layers are converted, pixel by pixel, to electronic image data which is then stored in memory circuitry for subsequent read-out and display on suitable electronic image display devices.
In typical digital radiography of the indirect type, a phosphor layer converts incident x-rays to visible light, which is then detected by a photosensor array that converts light intensity information to a corresponding electronic image signal. An intermediary fiber optic element may be used to channel the light from the phosphor layer to the photosensor array.
The perspective view of FIG. 1 shows a partial cutaway view of a small edge portion of a digital radiography (DR) detector 10 of the indirect type. A phosphor layer 12, formed from scintillating materials, responds to incident x-ray radiation by generating visible light that is, in turn, detected by a detector array 20. An optional fiber optic array can be provided for directing light from phosphor layer 12 toward detector array 20. Detector array 20 has a two-dimensional array having many thousands of radiation sensitive pixels 24 that are arranged in a matrix of rows and columns and are connected to a readout element 25. As shown at enlarged section E, each pixel 24 has one or more photosensors 22 and includes an associated switch element 26 of some type. To read out image information from the panel, each row of pixels 24 is selected sequentially and the corresponding pixel in each column is connected in its turn to a charge amplifier (not shown). The outputs of the charge amplifiers from each column are then applied to other circuitry that generates digitized image data that can then be stored and suitably processed as needed for subsequent storage and display.
Indirect DR imaging, using components arranged as in FIG. 1, shows promise for providing improved diagnostic imaging performance with high levels of image quality. However, some drawbacks remain. Because scintillating phosphor layer materials respond to incident x-ray radiation by emitting light over a broad range of angles, there is some inherent amount of scattering in the indirect detection process. Image sharpness is degraded when the visible light emitted from the phosphor is allowed to spread from its point of origin. The farther the emitted light spreads before detection by the photosensor, the greater the loss of light and sharpness. Any type of gap between the phosphor layer and its corresponding photodetector array can allow light to spread and consequent loss of image quality. For this reason, it can be particularly important to place the phosphor layer 12 (FIG. 1) as close to the photodetector (detector array 20) as possible.
In addition to losses from spreading and scattering, some further loss of light can occur due to reflection, such as where the light traverses an interface to a material with lower refractive index. Reflected light returning toward the phosphor layer may be reflected again by the phosphor and can travel to the photosensor in a position that is even farther from its point of origin, thus further degrading the sharpness of the image. This type of effect reduces the overall optical efficiency of image formation due to loss of light, signal crosstalk, and related effects, and tends to degrade image quality.
Phosphor layers used to convert x-rays to visible light in radiography are typically prepared by one of two methods. One method is to mix particles of phosphor with a binder and form this mixture into a sheet, usually by coating the mixture onto a carrier film. Another method is to evaporate phosphor onto a sheet substrate, forming needle-like structures. In both methods, the phosphor layer is covered with a protective coating to prevent physical and chemical damage.
The cross-sectional side view of FIG. 2 shows the layered arrangement of conventional digital detector 10 and shows where adhesive is commonly used. Phosphor layer 12 typically is provided on a substrate 14 and is optionally affixed to a fiber optic array 52, which is, in turn, affixed and optically coupled to detector array 20. An adhesive layer 28 is provided between detector array 20 and fiber optic array 52 and between fiber optic array 52 and phosphor layer 12. In conventional practice, substrate 14 may also support additional components as shown subsequently, including a carbon-pigmented black layer for absorbing leakage light and a pigmented white layer for reflecting some portion of the scattered light back through phosphor layer 12.
Among methods employed for improving optical coupling between the scintillator screen and the detector are the following, represented schematically in FIGS. 3A through 3F:
(i) Applying continuous pressure between the phosphor layer and the detector array, thereby maintaining physical contact between these assemblies. This type of solution, shown by arrows in FIG. 3A, can be difficult to maintain across the full surface of the detector. Moreover, it is difficult to make a digital radiography sensor as thin as necessary if mechanical clamping or hold-down devices are employed in order to maintain optical contact between the phosphor layer and photosensor array. Uniformity of optical contact is a must. Where an air gap occurs, the light transmission and the spatial resolution (MTF) would be significantly degraded.
(ii) Depositing the phosphor material directly onto the photodiode array of detector array 20. FIG. 3B shows a deposition apparatus 50 for forming scintillator layer 12. This method assures physical contact, hence good optical contact. However, this type of processing can be complex, may risk damage to the photodiode array and can be very expensive. Detector array 20 is an expensive device, making it impractical to use as a “substrate” for deposition or coating of materials. Uniformity of deposition also presents an obstacle that makes this type of solution less than desirable.
(iii) Use of a fiber-optic array 52, also termed a fiber optic plate or tile, between detector array 20 and phosphor layer 12, as shown in FIG. 3C. Array 52 is an optical device consisting of several thousands of glass optical fibers 54, each a few micrometers in diameter, bonded in parallel to one other. Each optical fiber acts as a light guide. Light from the radiation image is transmitted from phosphor layer 12 to the photodiode array of detector array 20 through each fiber 54. A typical fiber optic array is about 3 mm thick. Phosphor layer 12 is disposed on one surface of fiber optic array 52, then the other surface of fiber optic array 52 is pressed against detector array 20. The fiber optic array provides high-resolution imaging and, with some types of Complementary Metal-Oxide Semiconductor (CMOS) and Charge-Coupled Device (CCD) photosensor devices, can be useful for providing a measure of protection of the photosensors from high radiation levels. However, this is at the cost of considerable light loss (about 37%). Fiber optic array transmittance is about 63% for Lambertian light at the wavelength of 0.55 um. In addition, air gaps 44 can still occur on either surface of fiber optic array 52. This solution, therefore, also encounters the problems described in (i) and shown in FIG. 3A.
(iv) Depositing a phosphor layer directly onto the fiber-optic faceplate. FIG. 3D shows this hybrid solution. This solution reduces or eliminates air gaps 44 between phosphor layer 12 and fiber optic array 52; however, there can still be an air gap problem at the other surface of fiber optic array 52. This solution also suffers from lowered transmittance as at (iii).
(v) As in FIG. 3E, depositing a phosphor layer directly onto the fiber-optic faceplate as in (iv) and applying an optical adhesive 56 between the coated fiber optic array 52 and detector array 20. As with methods (iii) and (iv) just given, this method suffers from the inherently lower transmittance caused by the fiber-optic faceplate, fiber optic array 52.
(vi) As in FIG. 3F, insertion of a conventional optically transparent polymer layer 58 between phosphor layer 12 and detector array 20. The optical polymer materials used for this purpose may be in the form of fluid, gel, thermoplastic material, or glue. Each of these optical polymers has accompanying problems. Optical fluids are the most convenient to apply. However, as true fluids, they require containment or will otherwise tend to flow out from the optical interface if unsealed. Optical gels are non-migrating and do not require containment seals. However, they are too soft to provide dimensional rigidity, and may swell with prolonged exposure or at elevated temperatures. Optical thermoplastics (such as elastomers and resins) include soft plastics that, when cured, provide some dimensional rigidity. However, an additional thermal or radiation process for curing is generally required; such processing can be risky for electronic components of detector array 20. Optical glues exhibit similar problems as optical gels. It is also difficult to apply a uniform thickness of glue between the phosphor layer and the detector array. One solution for this problem, proposed in U.S. Pat. No. 5,506,409 to Yoshida et al. entitled “Radiation Detecting Device and the Manufacture Thereof”, is the use of spherical spacers to ensure the proper adhesive thickness. However, this requires a number of added steps for proper adhesion, with some complexity and risk of irregular spacer distribution.
Another method of constructing a digital radiography detector is to affix the phosphor layer directly to the fiber optic element or photosensor. In this case, there is an intervening layer of adhesive between the phosphor layer and the fiber optic array or photosensor. The phosphor may thus be optically coupled to the fiber optic element or to the photosensor, therefore reducing the amount of light that is reflected and refracted at the screen surface. This method is proposed, for example, in commonly assigned U.S. patent application Ser. No. 12/104,780 entitled DIGITAL RADIOGRAPHY PANEL WITH PRESSURE-SENSITIVE ADHESIVE FOR OPTICAL COUPLING BETWEEN SCINTILLATOR SCREEN AND DETECTOR AND METHOD OF MANUFACTURE by Yip, published as US 2009/0261259. To reduce the likelihood of losses due to reflection, the Yip disclosure proposes using an intermediary pressure-sensitive adhesive material between the phosphor layer and the photosensors and matching the refractive index of the pressure-sensitive material with that of the phosphor layer and that of the photosensor array. This method may provide a measure of improvement for rigid flat panel detectors that have relatively large imaging areas and can be advantageous where no fiber optic array is used. However, this method is not suited to the requirements of an image detector for dental imaging, where a low profile detector is most advantaged and where high image sharpness is a requirement. Use of intermediary materials in the light path can also be a disadvantage for applications in which more flexible detector materials are more desirable. Moreover, even where the index of refraction is closely matched to materials at the interface, any intervening adhesive layer increases the phosphor layer-to-detector distance over which the light tends to spread. Thus, sharpness degradation can still occur with this solution.
Thus, it is seen that there is a need for a digital radiographic detector that is suited for intra-oral imaging and that provides optical coupling between the photosensor array and the phosphor layer.